1. Technical Field
The subject matter disclosed herein generally relates to a low magnetic field and an ultra-low magnetic field nuclear magnetic resonance and magnetic resonance image apparatus. More specifically, the subject matter disclosed herein is directed to a low magnetic field and an ultra-low magnetic field nuclear magnetic resonance and magnetic resonance image apparatus using a prepolarization magnetic field coil of a type-II superconductor.
2. Related Art
Nuclear magnetic resonance (hereinafter referred to as “NMR”) is a phenomenon involved with precession of the magnetic spin of an atomic nucleus arising from resonance of the magnetic spin of the nucleus under a magnetic field when the magnetic field is applied to the atomic nucleus. Magnetic resonance imaging (hereinafter referred to as “MRI”) is a non-invasive technique of imaging the inner part of a target object by detecting electromagnetic waves generated during the precession using the NMR. The MRI is widely used as a medical diagnostics tool to image the inner part of human body.
Sensitivity of an MRI image is in proportional to to amount of magnetization and precession frequency. In general, very strong main magnetic field (main magnetic field; B0) using a superconducting magnet is applied to the target material to improve the sensitivity of the MRI image, since this strong magnetic field increases both the magnetization and the precession frequency.
Relaxation time of an NMR signal is in inverse proportion to uniformity of B0. Therefore, both magnitude and uniformity of B0 are important.
A superconducting magnet capable of generating a uniform magnetic field in the order of several Tesla (T) is very expensive. In addition, operation of the superconducting magnet requires liquid helium, which is an expensive refrigerant. Thus, the maintenance cost of an MRI system using the superconducting magnet increases.
Low/very low magnetic field NMR and MRI (hereinafter integrally referred to as “low magnetic field MRI”) is a novel MRI concept with an operating magnetic field in the order of microtesla to hundreds of microtesla. In a conventional MRI apparatus, both magnetization and precession frequency are increased, with strong B0, to improve sensitivity. However, in low magnetic field MRI, the requirement of main magnetic field is split into prepolarization magnetic field (Bp) and measurement magnetic field (Bm). The measurement magnetic field (Bm) may have a magnitude of several microtesla (uT) to tens of uT.
The low magnetic field MRI sequentially applies prepolarization magnetic field (Bp) and measurement magnetic field (Bm) to the sample. The prepolarization magnetic field (Bp) magnetizes the sample before being ramped down to zero. The prepolarization magnetic field (Bp) is significantly stronger than the measurement magnetic field (Bm) in order to magnetize the sample sufficiently. When the prepolarization magnetic field (Bp) is ramped down to zero, polarized nuclei spins precess around the measurement magnetic field (Bm). Thus, the precessing spins generate a time-varying magnetic field, which is then measured.
The prepolarization magnetic field (Bp) and the measurement magnetic field (Bm) are applied using separate coils independent of each other. The prepolarization magnetic field (Bp) is generated by a prepolarization magnetic field coil, and the measurement magnetic field (Bm) is generated by a measurement magnetic field coil. The prepolarization magnetic field (Bp) requires a strong magnetic field for sample magnetization, without the need for high field uniformity. The measurement magnetic field (Bm), on the other hand, requires a uniform magnetic field with low field magnitude. Thus, the low magnetic field MRI consists of a system with much simpler structure with lower manufacturing cost than conventional high magnetic field MRI using a superconductor main magnet.
With a low measurement magnetic field (Bm), a low-frequency spin precession in the order of tens of hertz (Hz) to hundreds of Hz. In conventional high magnetic field MRI, Faraday induction coils are used as receiver coils to measure the spin precession. Signal-to-noise ratio (SNR) of the Faraday induction coil is proportional to the measured signal frequency. For this reason, the Faraday induction is not suitable in measuring the low-frequency signal from the low magnetic field MRI. Thus, the low magnetic field MRI may use a superconducting quantum interference device (hereinafter referred to as “SQUID”) magnetic sensor that measures the time-varying magnetic field from the precessing spins directly to improve a low signal-to-noise ratio (SNR) of the Faraday induction coil, since the SQUID magnetic sensor has a flat frequency response characteristics.
Low magnetic field MRI may operate in a magnetic field in the order of microtesla using the SQUID. The low magnetic field MRI can image interior of an object from resonance signal with several to hundreds of Hz bandwidth, which is proportional to the magnitude of the measurement magnetic field (Bm). The low magnetic field MRI can significantly reduce distortion of an imagedue to magnetic artifact from surrounding metal. Thus, the low magnetic field MRI can observe a phenomenon that cannot be observed by a conventional high magnetic field MRI. In the conventional high magnetic field MRI, sample magnetization can be saturated and thus soft issue cancer cannot be identified without assistance of contrast agents. However, the low magnetic field MRI can image soft issue cancer without contrast agents.
The low magnetic field MRI can also be used reasonably on people wearing metallic prostheses or cardiac pacemakers. In addition, the low magnetic field can obtain images inside metal cans non-invasively. Thus, the low magnetic field MRI can be used as apparatuses supplementing X-ray devices widely used in security imaging.
The low magnetic field MRI may include a prepolarization magnetic field coil to magnetize a sample, a measurement magnetic field coil to induce nuclear precession from the magnetized sample, a SQUID magnetic sensor to read magnetic resonance signal from the nuclear precession, and a cooling system to cool the SQUID magnetic sensor to its operating temperature, which should be lower than the critical temperature of the superconductor comprising the sensor.
In low magnetic field MRI using a SQUID sensor, the prepolarization magnetic field coil may be made of superconductor. Superconductors are classified into two types, called Type-I and Type-II, based on their superconducting-to-normal transition when exposed to strong magnetic field. Lead (Pb) and tantalum (Ta), for example, are Type-I superconductors.
Conventionally, Type-II superconductors include metal alloys or oxide ceramics. Most high-temperature superconductors are Type-II superconductors. Niobium-titanium (NbTi), niobium-tin (Nb3Sn), and magnesium diboride (MgB2), for example, are metal alloy based Type-II superconductors.
There are some pure metal type-II superconductors like niobium (Nb), vanadium (V), and technetium (Tc). Oxide ceramic based Type-II superconductors include ReBCO (rare-earth-barium-copper-oxide) such as BSCCO and YBCO (yttrium-barium-copper-oxide).
For low magnetic field MRI, type-II superconductors are more economical, better in their physical properties, and have superior superconducting characteristics such as critical current and critical magnetic field than type-I superconductors.
A prepolarization magnetic field coil using superconductor is disclosed in Korean Patent Publication No. 10-2010-0076150. Since current density of a superconductor prepolarization magnetic field coil can be at least 100 times higher than that of a typical copper conductor coil, stronger magnetic field can be generated with less turns. In addition, the prepolarization magnetic field coil can be much smaller than a copper conductor coil. The superconductor prepolarization coil can operate at liquid helium temperature, whether the polarization coil conductor is a high-temperature superconductor or a low-temperature superconductor. Thus, a superconducting prepolarization coil can be integrated into the dewar containing the SQUID sensor.
None of known type-I superconductors have critical field strong enough to be used in a prepolarization magnetic field coil. For example, critical magnetic field of lead (Pb) is about 81 mT. Therefore, type-I superconductors are not suitable for prepolarization coils that generate a considerably high magnetic field in the order of tens to hundreds of mT, where self-field generated inside the conductor can be up to several times the desired magnetic field .
On the other hand, type-II superconductors have very high critical magnetic field. For example, critical magnetic field of NbTi is around 15 T. Unlike in type-I superconductors, however, t magnetic flux starts to penetrate the bulk of the superconductor at a relatively low magnetic field in type-II superconductors and the flux remain penetrating even when the magnetic field is removed, which is known as flux pinning phenomenon. This threshold field is known as the first critical magnetic field. The first critical magnetic field is lower than the critical field of a similar Type-I superconductor. Accordingly, when a type-II superconductor is used in a prepolarization magnetic field coil, the prepolarization magnetic field coil itself may be magnetized by the trapped magnetic flux from the strong prepolarization magnetic field and deteriorate resulting NMR signal. A method to remove trapped magnetic flux caused by superconducting magnetization hysteresis is required when the prepolarization magnetic field coil is made of type-II superconductor in order to prevent unwanted deterioration in NMR signal.